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Abstract:

A time-dependent decay behavior is incorporated into one or more joint
actuator control parameters during operation of a lower-extremity,
prosthetic, orthotic or exoskeleton device. These parameters may include
joint equilibrium joint impedance (e.g., stiffness, damping) and/or joint
torque components (e.g., gain, exponent). The decay behavior may be
exponential, linear, piecewise, or may conform to any other suitable
function. Embodiments presented herein are used in a control system that
emulates biological muscle-tendon reflex response providing for a natural
walking experience. Further, joint impedance may depend on an angular
rate of the joint. Such a relationship between angular rate and joint
impedance may assist a wearer in carrying out certain activities, such as
standing up and ascending a ladder.

Claims:

1. A prosthesis, orthosis or exoskeleton apparatus comprising: a proximal
member; a distal member; a joint connecting the proximal and distal
members, the joint adapted to permit flexion and extension between the
proximal and distal members; a motorized actuator configured to apply at
least one ofa joint impedance and a joint torque, the joint impedance
including at least one of a stiffness and damping, wherein the stiffness
is referenced to a joint equilibrium; a sensor configured to detect at
least one of a phase and a change in a phase of joint motion in a
repetitive cycle; and a controller configured to modulate at least one of
the joint equilibrium, the joint impedance and the joint torque, the
modulation employing a decaying time response as a function of at least
one of the phase and the detected change in phase of joint motion.

2. The apparatus of claim 1, wherein the sensor is configured to detect a
state transition phase of gait.

3. The apparatus of claim 1, wherein the apparatus is an ankle
prosthesis, orthosis or exoskeleton.

4. The apparatus of claim 1, wherein the stiffness is at least one of a
Swing-phase stiffness, a Controlled Plantar Flexion stiffness, a
Controlled Dorsiflexion stiffness and a Powered Plantar Flexion
stiffness.

6. The apparatus of claim 5, wherein the positive force-feedback
component comprises at least one of a gain and an exponent as applied to
at least one of the joint torque and the actuator torque.

7. The apparatus of claim 1, wherein the modulation is a function of at
least one of a proximal member angular rate, a distal member angular rate
and at least one of a joint torque rate and an actuator torque rate.

8. The apparatus of claim 6, wherein the gain and exponent are modulated
as a function of at least one of a proximal member angular rate, a distal
member angular rate and a torque rate.

9. The apparatus of claim 1, wherein the apparatus is a knee prosthesis,
orthosis or exoskeleton.

10. The apparatus of claim 1, wherein the stiffness comprises an early
stance flexion stiffness.

11. The apparatus of claim 1, wherein the stiffness comprises a knee
flexion stiffness that is a function of knee joint angular rate.

12. The apparatus of claim 1, wherein the joint torque is in a late
stance and is a positive force feedback component.

13. The apparatus of claim 5, wherein the positive force feedback
modulation is a function of the rate of change of the joint torque.

14. (canceled)

15. The apparatus of claim 6, wherein the gain and the exponent are
modulated according to at least one of the detected phase and a change in
the detected phase.

16. The apparatus of claim 1, wherein the decaying time response
comprises an exponential decay.

17. The apparatus claim 1, wherein the sensor is configured to detect a
joint position.

18. The apparatus of claim 1, wherein the controller is configured to
modulate the joint equilibrium to converge with the detected joint
position.

19. A method of controlling a joint impedance and a joint equilibrium of
a prosthesis, orthosis or exoskeleton apparatus, comprising: actuating a
joint of the apparatus; tracking a current joint position of the
apparatus; and controlling a value of the joint equilibrium of the
apparatus so as to converge to a value of the current joint position.

20. The method of claim 19, wherein converging a value of the joint
equilibrium of the apparatus comprises solving a differential equation
with adjustable coefficients.

21. A prosthesis, orthosis or exoskeleton device, comprising: a joint
constructed and arranged to permit flexion and extension between a
proximal member and a distal member; a motorized actuator configured to
apply at least one of a joint impedance and a joint torque, the joint
impedance referenced to a joint equilibrium; a sensor configured to
detect a characteristic of the device; and a controller configured to
modulate a parameter comprising at least one of the joint equilibrium,
the joint impedance and the joint torque according to the detected
characteristic, the modulated parameter exhibiting time-dependent decay
behavior.

22. The device of claim 21, wherein the joint torque comprises a positive
force-feedback component comprising at least one of a gain and an
exponent as applied to the joint torque.

23. The device of claim 22, wherein the positive force feedback component
comprises a function of a rate of change of at least one of a joint
torque and an actuator torque.

24. The device of claim 21, wherein the modulated parameter comprises at
least one of a proximal member angular rate, a distal member angular rate
and at least one of a joint torque rate and an actuator torque rate.

26. The device of claim 21, wherein the sensor is configured to detect a
joint position.

27. The device of claim 26, wherein the controller is configured to
modulate the joint equilibrium to converge with the detected joint
position.

28. A prosthesis, orthosis or exoskeleton device, comprising: a joint
constructed and arranged to permit flexion and extension between a
proximal member and a distal member; a motorized actuator configured to
apply at least one of a joint impedance and a joint torque, the joint
impedance referenced to a joint equilibrium; a sensor configured to
detect an angular rate of at least one of the proximal member, the distal
member and a joint connecting the proximal and distal members; and a
controller configured to modulate a parameter comprising at least one of
the joint equilibrium, the joint impedance and the joint torque according
to the detected angular rate to include at least one of a rate dependent
stiffness response and a decaying response.

30. The device of claim 28, wherein the joint torque comprises a positive
force-feedback component comprising at least one of a gain and an
exponent as applied to the joint torque.

31. The device of claim 30, wherein the positive force feedback component
comprises a function of a rate of change of the joint torque.

32. The device of claim 28, wherein the modulated parameter comprises at
least one of a proximal member angular rate, a distal member angular
rate, and a torque rate.

33. A prosthesis, orthosis or exoskeleton apparatus, comprising: a
proximal member; a distal member; a joint connecting the proximal and
distal members, the joint adapted to permit flexion and extension between
the proximal and distal members; a motorized actuator configured to apply
torque at the joint a sensor configured to detect at least one of a phase
and a change in a phase of joint motion in a repetitive cycle; a battery
to store electrical energy and to power the apparatus, a controller
configured to short the leads of the motor where the controller recovers
electrical energy from the apparatus during at least part of the
repetitive cycle.

34. The apparatus claim 33, wherein the energy recovery is accomplished
by pulse-width modulation of shorting the motor-leads.

35. The apparatus claim 34 wherein the pulse-width modulation is applied
to modulate an impedance response.

36. The apparatus claim 35, wherein the impedance has at least one of an
impedance and damping component.

37. The apparatus of claim 34, wherein the battery is not removable from
the apparatus.

38. A prosthesis, orthosis or exoskeleton apparatus comprising: a
proximal member; a distal member; a joint connecting the proximal and
distal members, the joint adapted to permit flexion and extension between
the proximal and distal members; a motorized actuator configured to apply
at least one of a joint impedance and a joint torque, the joint impedance
including at least one of a stiffness and damping, wherein the stiffness
is referenced to a joint equilibrium; a sensor configured to detect at
least one of a phase and a change in a phase of joint motion in a
repetitive cycle; and a controller configured to modulate at least one of
the joint equilibrium, the joint impedance and the joint torque, the
modulation response in the cycle determined by at least one of a joint
torque rate and an actuator torque rate during a part of the cycle.

Description:

BACKGROUND

[0001] 1. Field of the Invention

[0002] Devices and control systems for biologically-inspired artificial
limbs are generally disclosed.

[0003] 2. Related Art

[0004] Existing prosthetic leg devices include a series-elastic actuator
which functions as a biologically-inspired muscle-tendon unit to
modulate, during a gait cycle, joint impedance, joint equilibrium and
torque, in accordance with walking speed and terrain modality (e.g.,
sloping ground, stairs, etc.). It is desired for prosthetic leg devices
to function in a way that matches the human ankle response as captured,
in part, by FIG. 1, which illustrates human biomechanical function in a
gait cycle, on level-ground. In the schematic of FIG. 1, the gait cycle
on level-ground is initiated by a heel-strike event. Other types of gait
cycles, such as toe-strike initiated cycles as might occur in steep ramp
or stair ascent, are not expressly shown.

[0005] Prosthetic leg devices have been designed so as to exhibit response
behavior captured by a "dashboard" of biomechanical characteristics,
shown in FIG. 2a. These biomechanical characteristics are based on
body-mass normalized and walking-speed reference measures from an intact
ankle population, including Net Non-Conservative Work, Peak Power,
Toe-off Angle and Peak Power Timing. As depicted in FIG. 2a, dashed lines
denote +/- sigma error bounds for the normative data, solid lines denote
average values for the normative data, and circles represent individual
step data wirelessly acquired from an ankle device wearer.

[0006] The ankle device depicted in FIG. 2b employs a state machine,
implemented in the intrinsic control firmware of the device to modulate
the actuator response. The actuator response is programmed to define a
joint impedance, joint equilibrium and torque, so as to emulate human
function in each gait cycle state. Depending on the phase of gait, the
device will enter into an appropriate state. At times, the transition(s)
between states for an artificial leg device may be abrupt, or might not
accommodate for changes in wearer intent.

SUMMARY

[0007] The inventors have recognized and appreciated there to be
advantages in employing time-dependent decay behavior in one or more
control parameters when the actuator torque of an artificial leg device
is modulated during use. While not meant to be limiting, such parameters
may include joint equilibrium, joint impedance (e.g., stiffness, damping)
and/or joint torque components (e.g,. gain, exponent) of the programmable
state (e.g., powered reflex response). The decay behavior may conform to
any suitable mathematical relationship, such as an exponential decay,
linear drop, quadratic function, piecewise relation, dynamic behavior
model that might arise from the output of a linear or non-linear
differential equation, or other suitable function. Such behavior, when
used in a positive force feedback system, may provide for a smooth
experience that emulates biological kinetics (torque, power) and
kinematics. For example, this type of control may ease the transition(s)
between states of the device (e.g., so that they are generally
unnoticeable to the wearer) and may allow for the wearer to alter his/her
course during gait in a natural manner.

[0008] In an illustrative embodiment, a prosthesis, orthosis or
exoskeleton apparatus is provided. The apparatus includes a proximal
member; a distal member; a joint connecting the proximal and distal
members, the joint adapted to permit flexion and extension between the
proximal and distal members; a motorized actuator configured to apply at
least one of a joint impedance and a joint torque, the joint impedance
including at least one of a stiffness and damping, wherein the stiffness
is referenced to a joint equilibrium; a sensor configured to detect at
least one of a phase and a change in a phase of joint motion in a
repetitive cycle; and a controller configured to modulate at least one of
the joint equilibrium, the joint impedance and the joint torque, the
modulation employing a decaying time response as a function of at least
one of the phase and the detected change in phase of joint motion.

[0009] In another illustrative embodiment, a method of controlling a joint
impedance and a joint equilibrium of a prosthesis, orthosis or
exoskeleton apparatus is provided. The method includes actuating a joint
of the apparatus; tracking a current joint position of the apparatus; and
controlling a value of the joint equilibrium of the apparatus so as to
converge to a value of the current joint position.

[0010] In yet another illustrative embodiment, a prosthesis, orthosis or
exoskeleton device is provided. The device includes a joint constructed
and arranged to permit flexion and extension between a proximal member
and a distal member; a motorized actuator configured to apply at least
one of a joint impedance and a joint torque, the joint impedance
referenced to a joint equilibrium; a sensor configured to detect a
characteristic of the device; and a controller configured to modulate at
least one of the joint equilibrium, the joint impedance and the joint
torque according to the detected characteristic, the modulation
exhibiting time-dependent decay behavior.

[0011] In a further illustrative embodiment, a prosthesis, orthosis or
exoskeleton device is provided. The device includes a joint constructed
and arranged to permit flexion and extension between a proximal member
and a distal member; a motorized actuator configured to apply at least
one of a joint impedance and a joint torque, the joint impedance
referenced to a joint equilibrium; a sensor configured to detect an
angular rate of at least one of the proximal member, the distal member
and a joint connecting the proximal and distal members; and a controller
configured to modulate a parameter comprising at least one of the joint
equilibrium, the joint impedance and the joint torque according to the
detected angular rate to include at least one of a rate dependent
stiffness response and a decaying response.

[0012] In yet another illustrative embodiment, a prosthesis, orthosis or
exoskeleton apparatus is provided. The apparatus includes a proximal
member; a distal member; a joint connecting the proximal and distal
members, the joint adapted to permit flexion and extension between the
proximal and distal members; a motorized actuator configured to apply
torque at the joint; a sensor configured to detect at least one of a
phase and a change in a phase of joint motion in a repetitive cycle; a
battery to store electrical energy and to power the apparatus, a
controller configured to short the leads of the motor where the
controller recovers electrical energy from the apparatus during at least
part of the repetitive cycle.

[0013] Other advantages and novel features of the invention will become
apparent from the following detailed description of various non-limiting
embodiments when considered in conjunction with the accompanying figures
and claims.

BRIEF DESCRIPTION OF THE DRAWINGS

[0014] Aspects of the present disclosure are described with reference to
the following drawings in which numerals reference like elements, and
wherein:

[0047] Various embodiments of the present disclosure relate to a
biologically-inspired, sensing and control architecture for bionic leg
actuation (e.g., knee joint actuation, ankle joint actuation). As
described herein, a bionic device may function to restore or replace
anatomical structure(s) and/or exhibit physiological process(es), with
one or more electro-mechanical components. For instance, bionic devices
of the present disclosure may emulate stance-phase kinetics (e.g., torque
and power) that may occur naturally in intact limbs. Bionic leg joints
described herein may employ a series-elastic actuator (SEA) to amplify
mechanical power, to enable closed-loop torque control and to enable
sensing of actuator torque through a model of the torque-displacement
characteristics. In some embodiments, an ankle device may employ a
hardstop with known flexion characteristics that limits dorsiflexion
travel of the joint. A control system modulates joint impedance (e.g.,
stiffness, damping), joint equilibrium (e.g., equilibrium location) and
joint torque (e.g., motor reflex gain, motor reflex exponent) in
accordance with gait cycle state and walking speed, a surrogate for
walking speed, or the rate of change of a state variable or sensor in the
actuator control system. In some embodiments, the rate of change of the
state variable may include an inertial pitch rate (e.g., of a tibial
component) and/or an actuator torque rate (e.g., of an ankle or knee
joint), shortly after foot strike.

[0048] In some embodiments, one or more parameters controlled by the
system may exhibit time-dependent behavior. For example, the joint
impedance, joint stiffness, joint damping, joint equilibrium, reflex
torque gain, reflex torque exponent, or another suitable parameter(s) may
employ a time decay (e.g., value of the parameter diminishes over time)
during an appropriate phase of gait. Such a decay may exhibit any
suitable functional behavior, such as exponential, linear, piecewise,
etc. This type of behavior, in some cases, may also provide for a natural
experience to the wearer, for example, without producing a feeling of
abruptness upon changes in the phase of gait. For instance, a gradual
lessening of ankle stiffness upon entry into an Early Stance mode may
allow for a wearer to rollover smoothly in a natural manner such that
mode changes (i.e., state transitions) of the device are transparent
(e.g., almost unnoticeable).

[0049] As used herein, a phase of gait may describe a particular state of
the device, which may be triggered by a gait event (e.g., heel-strike,
toe-off). For example, a phase of gait may refer to: a state transition
in a leg prosthesis control system, such as in a joint actuator
controller; the inertial state of proximal and distal members of the
device; and/or changes in one or more components of the inertial state of
the proximal and distal members of the device.

[0050] As used herein, a motorized actuator or motorized actuation system
may include any suitable motor. For example, motorized actuators may
incorporate one or more electric motors, hydraulic motors, pneumatic
motors, piezo-actuated motors, shape-memory motors, electro-polymer
motors, or any other appropriate motorized device.

[0051] As used herein, a characteristic of motion of a device may include
one or more of the following: an inertial pose of distal and proximal
members of the device; changes in the inertial pose of the distal and
proximal members of the device; translational velocity or angular rate of
one or more points on the distal and proximal members; kinetics,
including force, torque and power, and the derivatives thereof at the
joints and at the interface between the device and ground; kinematics,
including joint angles, and derivatives thereof; dynamic actuator
state(s), including force, torque, displacement in the motor drive and
transmission, including the elastic elements embodied within the
transmission; and other appropriate characteristics.

[0052] While neuroscientists identify increasingly complex neural circuits
that control animal and human gait, biomechanists have found that
locomotion requires little outside control if principles of legged
mechanics are heeded that shape and exploit the dynamics of legged
systems. Embodiments according to the present disclosure may include
muscle reflex response(s) that encode principles of legged mechanics, and
provide a link to the above observations surrounding the behavior of
natural limbs. Equipped with reflex control, various embodiments of
bionic devices presented herein reproduce human walking dynamics and leg
kinetics and kinematics; tolerate ground disturbances; and adapt to
slopes without outside parameter intervention(s), such as might otherwise
be informed by inertial sensor inputs, neural or cognitive functions.
Accordingly, aspects/parameters of the bionic response may be
appropriately encoded to adaptively modulate one or more parameters based
upon intrinsic kinematic and kinetic measures (e.g., angle and torque
including their derivatives) or extrinsic interventions arising from
measures of walking speed and terrain (as might be supplied by an
inertial measurement unit, for instance), so as to suitably emulate the
muscle-tendon reflex. Aspects described herein may employ principles
described in the article by Geyer, H. and Herr, H., entitled "A
Muscle-Reflex Model that Encodes Principles of Legged Mechanics Produces
Human Walking Dynamics and Muscle Activities," submitted to IEEE
Transactions on Neural Systems and Rehabilitation Engineering and
accepted in 2010, the disclosure of which is hereby incorporated herein
by reference in its entirety.

[0053] It can be appreciated that embodiments of the present disclosure
are not required to incorporate a state machine that transitions from one
discrete state to another in a gait cycle. For instance, a mere change in
inertial state across a gait cycle (e.g., based on the use of a rate
gyroscope to measure a rate of tibial pitch) may be a part of a gait
cycle phase.

[0054] Systems described herein may be incorporated in devices made by
iWalk, Inc., such as in the BiOMT2. In some cases, the BiOMT2
device employs a series-elastic actuator (SEA) that incorporates a
biophysically-based, reflexive control system. This system emulates
dominant muscle-tendon behavior, during walking, of the ankle plantar
flexors, the Soleus and Gastrocnemius calf muscles, as well as the
dominant dorsiflexor, the Tibialis Anterior. The SEA may control ankle
joint impedance (e.g., stiffness, damping), virtual spring equilibrium
and/or reflexive torque. The SEA system may enable sensing of actuator
torque (ΓSEA) through measurements of series-spring
deformation. Additionally, the ankle joint may include a hardstop, which
limits the ability for the ankle to move to a position of increased
dorsiflexion, after a certain point. In addition to measuring actuator
torque, the system may also monitor hardstop torque (Γhs)
through the measurement of hardstop spring deformation.

[0055] A finite state machine may be employed in a State Control Processor
to control transitions of the device through different states. The gait
cycle states in the State Machine may include early stance, late stance,
late stance power, early swing and late swing, which are aligned with the
conventional names employed in human biomechanics, namely, controlled
plantar flexion, controlled dorsiflexion, powered plantar flexion, early
swing and late swing, respectively. The transitions between these walking
gait phases may be determined by a system clock (time) and/or the SEA
torque (ΓSEA), hardstop torque (Γhs), and their
time derivatives.

[0056] In some embodiments, the device includes a single finite state
machine for walking. As a result, when a single finite state machine is
employed, the control system does not revert to a non-walking state
machine based on biomechanical change(s) made by the human wearer.
Accordingly, the device is less cumbersome than would otherwise be the
case if multiple state machines are incorporated.

[0057] The system may make some or all motor control actuation decisions
based upon kinetic sensory information of the device (e.g., force/torque
information), without requiring kinematic sensory information of the
device (e.g., positions, velocities, accelerations). For example, the
system is not required to employ reflex response parameter interventions
as these might be informed by accelerometers or rate gyros or any other
sensor for the measurement of overall device positions, velocities or
accelerations relative to horizontal or vertical reference planes to
adapt to walking speed and terrain modality. As a result, the position of
the ankle joint may be controlled based on the interaction forces
experienced between the human wearer, the device, and the ground surface.
Therefore, contrary to conventional robotic systems, it is not necessary
for the device to directly control the position of the ankle joint,
whether in stance or swing phases, as systems described herein are
controlled based on reflex response(s). Though, it can be appreciated
that, in some cases, the system may employ position sensors,
accelerometers, rate gyros and/or any other sensor, as suitably desired.

[0058] Non-linear, positive force feedback control is applied in powered
plantar flexion to emulate human muscle-tendon reflex dynamics. Devices
described herein employ positive force feedback with intent to emulate a
natural, uncontrolled (e.g., automatic) reflex response. This reflex is
implemented by a motor torque control that behaves according to a
positive force feedback mathematical relationship involving parameters
that include torque gain and torque exponent, each modulated according to
the stimulation of certain parameters, for example, the torque rate
measured by a series elastic actuator and/or the torque measured at a
hardstop.

[0059] The system control architecture employs motor and joint angle
sensing to compute, via calibrated models, instantaneous SEA and hardstop
torque. Instead of using inertial information, the system architecture
employs intrinsic measures of torque, torque rate of change and time
duration within a gait cycle state to inform transitions in the State
Machine that directs the response modulation in a Motor Processor and, in
some embodiments, may rely exclusively on torque and time within a state
to inform the transitions. That is, measurements of inertial information,
such as position, velocity and acceleration are not used to inform
parameter interventions that modulate the actuator response. Rather,
force measurements, such as force and torque measured over time, may be
used as input to direct the response modulation of the joint actuator.

[0060] The device may exhibit reflexive behavior, without any system
memory. That is, the system may monitor device torque(s) and reflexively
respond to such torque(s) with little delay between sensing and
actuation. As a result, the monitoring of torque throughout or during a
portion of a gait cycle may be the basis for modulation of control
actions during a current gait cycle, without any consequence to control
actions that affect a subsequent gait cycle.

[0061] In some embodiments, the control system does not require detection
of particular gait patterns or events, and in response, the control
system is not required to modulate either the control algorithm, or its
system parameters. The control algorithm and its parameters are not
necessarily adjusted in any manner in response to a user transitioning
from a walk to a run, nor while ambulating from a level-ground surface to
an incline, nor from level-ground to steps, nor while moving to standing,
nor from a standing position to a sitting position, nor from a standing
position to a leaning position, nor from a sitting position to a lying
down position, nor while putting on pants. That is, despite the type of
action the wearer may currently be performing, the control system may
function according to a single state machine control, without regard to
the type of user action currently performed.

[0062] The control system may be configured to detect a foot strike with
the ground surface based on torque/force information. Independently of
how the device has struck the ground, whether it is a heel strike, a toe
strike, or a foot-flat strike, the system may run the same algorithm with
the same control parameters.

[0063] Further, walking speed may be estimated from a known linearly
correlated relationship with normalized, peak derivative of SEA torque in
late stance. That is, torque rate may be used as an estimate (or
surrogate) of a current walking speed so as to inform the reflex
parameter modulation. In particular, the gain and exponent parameters of
a reflex relationship may be modulated based on a rate of change of a
parameter (e.g., pitch rate, torque rate). For example, a rate-based
blending (interpolation) of the parameters may be employed.

[0064] In addition, to achieve a smooth and natural response, in some
embodiments, the stiffness and/or damping of the joint in Early Stance
may be designed to decay exponentially, for example, smoothly reducing
stiffness/damping so as to increase joint compliance. Such exponential
decay behavior, for impedance, may be particularly beneficial for a
wearer of an artificial leg device when walking slowly on uneven terrain
or descending down a steep slope, allowing for seamless, hi-fidelity
device control.

[0065] In some embodiments, artificial leg devices are constructed
according to a biologically-inspired approach where an IMU is not
required for their use. A number of design principles are considered in
constructing the artificial leg device.

[0066] For example, the time duration in a state, torque and torque
derivative (torque rate) may guide the device in transitioning from one
state to another, as well as to modulate the reflex parameters, which may
or may not correlate with a current walking speed. In some cases, a
single measured parameter may be sufficient as a signal for transitioning
the device between states and/or estimate walking speed. As discussed,
time duration within a state, SEA torque (ΓSEA) and hardstop
torque (Fhs)--and the time derivatives of these--may be used as
parameters that the system uses to inform state transitions and, in some
cases, may be used independently and/or exclusively from other
parameters. Peak SEA torque rate as sampled during late stance may be
employed in the adjustment of the late stance power reflex, which may
occur independently of an estimation (or correlation) of walking speed.
As such, it may be a useful observation, yet not necessary for
embodiments of the present disclosure, that the above-mentioned rate(s)
may correlate with walking speed, for a broad range of wearers. As such,
it is not necessary in the preferred embodiment to explicitly estimate
the walking speed and to use that estimate to inform the reflex response
modulation. So, the intrinsic inertial, kinematic or kinetic may be used
directly to inform that modulation.

[0067] As muscle-tendon units of an intact limb do not employ inertial
sensing to modulate their response, such intrinsic measures may enable
the device to behave and respond as a more natural muscle-tendon unit.
Instead, in an intact ankle, muscle and tendon stretch (torque) and their
various rates of change are key inputs to the spinal reflex arc
connecting the tendon and the muscle. As a result, transitions are more
natural and consistent even when the wearer walks softly or runs and
jumps in place.

[0068] Further, the system may employ a uniformly-applied
stiffness/impedance that decays smoothly after foot strike. When the
impedance after foot strike is set to decay, "impedance switching"
between states, and the abrupt nature that often accompanies such a
switch, may be eliminated. Early Stance impedance--generally defined by
stiffness (kes) and damping (bes)--may be used by all states,
except, in some cases, it might not be used during late stance power and
early swing. Impedance may be set in late-swing to a programmable (tuned)
value. In some embodiments, kes decays exponentially to a
programmable value, kes.sub.∞, which is typically a small
fraction of the initial value, kes0.

[0069] Exponential decay of impedance, or one or more other appropriate
parameters, may begin at entry into Early Stance. In some cases, the time
constant for decay may be set so that the stiffness is substantially
maintained (e.g., does not drop quickly) during controlled plantar
flexion (CP) (e.g., a time duration between 0.05-0.2 seconds), such as
when walking at a brisk walking speed. When walking more slowly, e.g.,
down a steep hill, the stiffness may be set to drop smoothly, or more
quickly, so as to enable the foot to find an equilibrium state at
foot-flat with a diminished spring restoring torque--thereby reducing
socket stress. The exponential decay behavior (e.g., for joint impedance,
joint equilibrium, torque, or others) may continue for a portion of or
for the entire gait cycle. For instance, in some cases, exponential decay
may continue until it is reset at entry into Early Stance. Such
transitions may occur without the wearer even noticing the occurrence of
a state transition--thereby eliminating confusion and irritation.

[0070] A single walking state machine may deliver a biomimetic response
either while walking or not walking, without need for a secondary
non-walking state machine. Instead of discretely switching between a
non-walking state machine and a walking state machine, state machines of
the present disclosure may use the Early Stance state to uniformly
deliver a biomimetic response without having to reconfigure the joint
impedance and/or joint equilibrium when in a non-walking state. To
accomplish this, the walking state machine may cause transition(s) to
Early Stance if the time duration within any of the other walking machine
states exceeds a programmable limit for that state, typically about two
seconds. The stiffness, kes, may continue to decay to deliver a
smoothly varying impedance that, in the limit, devolves to a
substantially lightly damped response that responds naturally for
non-directed activities that do not involve locomotion. As discussed
above, for some embodiments, only torque and torque derivatives are used
to inform the logic transition between states, for example, from early
stance to late stance and late stance power where locomotion may then be
initiated.

[0071] In some embodiments, spring impedance (e.g., stiffness, damping)
may be dependent on angular rate in, for example, an ankle or a knee. For
instance, an artificial joint device may employ a bionic control system
that modulates the impedance of the joint so as to assist the wearer
during stair ascent, steep ramp ascent or during the transition from
sitting to standing. In some cases, when flexed past a certain threshold
angle, the spring stiffness of the joint may be rate dependent, applying
positive feedback in response to increases in the joint angular rate or
the absolute value of joint angular rate. As an example, the spring
stiffness of an artificial knee joint may be modulated such that when a
wearer is standing up and the angular rate is increased, the joint
becomes stiffer so as to provide increased support during the standing
motion. Such support is effective to assist the wearer in standing up.

[0073] In particular, concepts described herein may be guided by design
principles that motivate use of positive force feedback, use of
intrinsic, motor damping behavior to implement dynamic clutches, and
catapult behaviors, such as those described in U.S. patent applications
entitled "Variable-Mechanical-Impedance Artificial Legs" corresponding to
Ser. Nos. 60/395,938; 10/613,499; 13/363,820, the disclosures of each of
which are also hereby incorporated herein by reference in their entirety.

[0074] It should be understood that for those skilled in the art, the
control architecture described herein may be extended to bionic ankles
that employ physical and/or SEA-applied virtual, unidirectional and
bi-directional parallel elastic elements where torque-displacement
characteristics of these systems may be calibrated before use. Further,
while such control architecture(s) may be applied to a bionic ankle
prosthesis, these principles may be readily extended to orthotic,
exoskeletal or humanoid applications in lower-extremity augmentation of
ankle, knee and hip.

[0075] While systems in accordance with the present disclosure do not
require inertial measurements as input for actuator modulation, it can be
appreciated that systems described herein may be used in place of or in
combination with inertial measurement systems. For instance, an actuator
response may be accomplished by controlling motor torque, τm, in
a closed-loop or open-loop manner, to match a desired response. In such
an architecture, joint angle, motor angle and 6-DOF inertial state
(orthogonally-opposed measures of local angular rate and acceleration as
sampled by an Inertial Measurement Unit (IMU)) may be used to compute SEA
and hardstop torque via calibrated models, to inform state machine
transitions, to estimate walking speed and/or to adapt to changes in
walking speed or terrain modality. As discussed above, SEA torque and
hardstop torque may be used as input to modulate reflex parameters
employed in powered plantar flexion. Table 1 provides a summarized
mapping of the intrinsic firmware states to the level-ground, gait cycle
states as implemented in an artificial ankle device. FIG. 3 shows a
schematic of an artificial ankle device that illustrates various
parameters that may be referenced in the present disclosure.

[0076] In systems that operate under the firmware states summarized by
Table 1, the State Machine employs state transitions that are informed by
time duration within the state, actuator torque, hardstop torque, and
inputs from the Inertial Measurement Unit (IMU). Complex measures of
"jerk" and vibration applied to the z-component of the local or
world-referenced acceleration are employed to detect heel or toe strike
transition from late swing (LSW) to early stance (ES). Logic employing
pitch velocity (tibia rotation in the sagittal plane) is used as a
"guard" (qualifying) condition prior to applying the accelerometer-based
foot strike logic. Pitch velocity, as measured at or near the entry into
late stance (LS) may be used (as a surrogate) to estimate walking speed
and as input for determining resulting reflex response parameters
(pff({dot over (s)}) and N({dot over (s)})) in late stance power
(LSP).

[0077] Further, pitch rate or velocity may be used to inform state
transitions from a non-walking state machine into a walking state
machine. While such an IMU-based approach may work well for normal gait
cycles involving locomotion (e.g., walking), such an approach might not
be optimized for non-walking type sequences, for example, those that may
occur when the wearer is moving slowly in a confined space, moving
between standing and sitting positions, or ascending/descending a ladder.
In a small percentage of such cases, a completely IMU-based actuator may
have a tendency to respond more vigorously than desired. Conversely, in
situations where the wearer is running or jumping in place, the state
machine might miss an occasional transition, thereby causing the actuator
response to be, in some cases, inconsistent.

[0078] The impedance response when the system is set to a non-walking
state may, at times, be constrained to be a viscous damper (e.g., have a
high damping coefficient resulting from shorting of the motor leads) for
a discrete period of time (e.g., approximately two seconds) followed by a
more lightly-damped response, which is a less than natural response for
the wearer. In cases where transitions between non-walking and walking
occur over short time intervals, the step response in viscosity may
become less than desirable.

[0079] Considering again artificial leg devices that are programmed in a
biologically-inspired manner where an IMU is not required, Table 2
provides a summary for such a device. Such devices may be constructed and
programmed to capture the reliance on torque-time and the use of an
exponential decay so as to eliminate or reduce the abruptness that may
result due to transition from one state to another.

[0080] Table 2. Alignment of level-ground gait cycle states with intrinsic
firmware states for an embodiment.

[0081] To those skilled in the art it should be readily apparent that the
computation and prediction of walking speed is not necessary. In some
embodiments, the reflex parameters can be computed as a function directly
of the SEA torque rate without loss of generality in another preferred
embodiment. FIG. 4 illustrates various state transitions that may occur
throughout two typical gait cycles--first exiting from Early Stance into
two successive heel-strike first gait cycles. Note that for convenience,
virtual state 1 is used as a representation of Early Stance at t=∞.
As described earlier, the early stance stiffness, kes, decays
exponentially leaving the damping, bes, as the dominant impedance
component.

Early Swing (ESW) to Late Swing (LSW) Transition

[0082] As shown in FIG. 4, the ESW-LSW (2-3) transition may occur at a
fixed time (e.g., approximately 100 msec, between about 10 msec and about
200 msec) after entry into ESW. During ESW, an overdamped, second-order,
joint equilibrium trajectory is launched, that returns the ankle angle,
θ, back to θes--a position at or near the neutral
position so as to avoid a tripping hazard. In some embodiments, the time
constant, τesw applied in this trajectory is between about 10
msec and about 150 msec (e.g., approximately 50 msec), so as to
correspond with that of an intact human ankle.

Late Swing (LSW) to Early Stance (ES) Transition

[0083] FIG. 5 illustrates an embodiment of a state transition from Late
Swing to Early Stance (3-4). The embodiment shows the hardstop
(Γhs) and SEA (torque ΓSEA, torque rate {dot over
(Γ)}SEA) torque component response for a heel-strike, first
transition. The state and motor ready flags are also shown. In this
example, the motor ready flag denotes the motor controller state. As
shown in this figure, a value of -2 denotes that an ankle trajectory is
running and has not yet finished. A value of +2 denotes that the ankle
trajectory has completed and that a motor coil resistance measurement is
being acquired. A value of 6 denotes that the motor controller is ready
to apply animpedance or respond to a new trajectory or function command.

[0084] FIG. 6 depicts another embodiment of a state transition from Late
Swing to Early Stance (3-4). Instead of a heel-strike transition, this
embodiment shows a toe-strike transition. As can be seen, a substantial
difference between the two different ground impact conditions is that in
the situation where heel-strike occurs first, the ground impact imparts a
large negative torque, ΓSEA, and a large negative torque rate
{dot over (Γ)}SEA, on the SEA. Whereas in the case where
toe-strike occurs first, the ground impact imparts a large positive
torque, Fhs, against the hardstop. As such, to detect these
conditions reliably, a "guard condition" may first be applied to the
state transition logic so as to reject the "noise" in ΓSEA and
Fhs, during the swing phase--this is a result of the SEA torque
applied to achieve the ankle trajectory and a possible collision with the
hardstop during the time interval.

[0085] Accordingly, for each type of state transition, a threshold would
be crossed (e.g., when the measured or sensed torque is greater than or
less than a particular set torque value, within a certain period of time)
that triggers transition from one state to

Walking-Speed Referenced Reflex

[0086] The device may use the maximum, rate-of-change in SEA torque ({dot
over (Γ)}SEA) as measured in Late Stance as an estimation (or
surrogate) for instantaneous walking speed. FIG. 10a illustrates data
that shows a linear relationship that exists between {dot over
(Γ)}SEA and the tibia pitch rate, Ψ. The tibia pitch rate
at mid-stance (after the foot flat condition) is further known, through
experimentation, to be proportional to leg-length normalized walking
speed, as shown in FIG. 10b and as discussed in U.S. patent application
Ser. No. 13/079,564. This estimation of walking speed may be computed
just before use in Late Stance Power to inform the reflex parameter
modulation.

[0087] The graph shown in FIG. 10c reports a high degree of correlation
(R2) of pitch velocity vs. SEA torque rate during Late Stance that
exists across a broad population of production units and walkers (see
circles), as measured in a standard walkabout test used to create a
Dashboard.

[0088] Such studies have shown that {dot over ({tilde over
(Γ)}SEA is not invariant across a population of wearers, even
when normalized by, for example, peak torque at a self-selected walking
speed. So, in one embodiment, {dot over (Γ)}SEA is observed
for each specific wearer--both at the fastest achievable walking speed
and at the slowest desired walking speed. At each speed, preferred values
for torque gain, pff ({dot over (s)}), and torque exponent, N({dot
over (s)}), may be determined by tuning--thereby determining
values/ranges for various parameters, such as pff slow Nslow,
Pff fast, Nfast. With these parameters in hand, a basis is
provided through which the reflex response may be blended across a range
of walking speeds. By replacing {dot over (s)} with {dot over
(Γ)}SEA, the following blended reflex equations may be used:

[0092] Where the subscript, SEA, on {dot over (Γ)}SEA, is
removed to simplify the notation.

Device Extensions

[0093] It should be appreciated that while device control architectures in
accordance with the present disclosure have been applied to an artificial
(bionic) ankle device with a hardstop, the hardstop functionality may be
replaced by a physical, unidirectional or bi-directional element,
parallel elastic element, a virtual, SEA-applied, parallel elastic
element, or other suitable component. For example, in either case the
hard stop torque, Γhs, may be replaced by a parallel elastic
element torque, ΓPE, where ΓPE is calibrated in
manufacturing to determine the torque displacement characteristics of the
physical or virtual elasticity.

[0094] Further, while device control architectures described herein have
been applied to artificial ankle prostheses, concepts presented here may
be extended for application in orthotic, exoskeletal, humanoid ankles, or
other appropriate devices. And, while the device control architectures
herein have been applied to artificial ankle applications, the techniques
applied here may also be extended for use in accordance with other
lower-extremity applications, for example, in the knee and hip.

Further Embodiments and their Implementation for Prosthetic or Orthotic
Ankle Devices

[0095] Embodiments of bionic leg devices, such as the BiOMT2 system
produced by iWalk, Inc., may employ five states--Early Stance (ES; State
4), Late Stance (LS; State 5), Late Stance Power (LSP; State 6), Early
Swing (ESW; State 2) and Late Swing (LSW; State 3)--that align with the
human biomechanical gait cycle states controlled plantar flexion (CP),
controlled dorsiflexion (CD), powered plantar flexion (PP), Early Swing
(ESW) and Late Swing (LSW), respectively. The present disclosure reviews
various details of control actions within each state and describes the
state transition logic that causes entry into the state.

[0096] Early Stance (ES) Control Action

[0097] In ES (State 4), for some embodiments, the SEA applies a
lightly-damped, torsional spring response in accordance with the human
biomechanical joint response in Controlled Plantar Flexion. The impedance
as applied by the SEA motor torque, τm, is comprised of a
time-varying spring, kes(t), and a time-varying damping component,
bes(t). The "virtual spring" joint equilibrium, θes, is
the ankle angle as captured at ES entry. In some cases, one or more
variables (e.g., spring constant, damping component, joint equilibrium,
gain, exponent, etc.) of the motor torque may be time-dependent and/or
may exhibit a time decay-type behavior (e.g., exponential, linear,
piecewise, etc.). The actuator may apply an exponential decay to the
stiffness component in order to make the ankle increasingly more
compliant as the state progresses--to emulate human biomechanics while
walking slowly, including on steep or uneven terrain. The ES control
action may be modeled as follows:

where τm is the motor torque, θ is the joint angle, β
is the SEA motor angle, And where, τes{dot over
(k)}es(t)+kes(t)=kes.sub.∞ applies an exponential
stiffness decay with time constant, τes
θes=θ(t=0), In some embodiments, the following
second-order relation may be used to model exponential stiffness decay:

τkes2{umlaut over
(k)}es(t)+2τkes{dot over
(k)}es(t)+kes(t)=kes.sub.∞

t=time since ES entry kes (0)=kes.sub.∞,
bes(0)=bes0 To those skilled in the art, other linear or
non-linear differential equations can be applied to accomplish this decay
function. As provided in the equation above, the stiffness decays to
kes.sub.∞ with a time constant, τes--e.g., about 200
milliseconds, or between 100-500 milliseconds. In some embodiments, the
time constant may be set (e.g., optimized) so as to allow the ankle to
conform to the ground surface while the wearer walks slowly down an
incline. Examples of these are included in Table 3 below.

Early Stance (ES) Entry State-Transition Details

Late Swing (LSW)-to-Early Stance (ES) Transition

[0098] In some embodiments, the state transition into ES from LSW may
occur when a foot-strike is detected--for example, by presence of a large
or increasing heel load (L3-4B or L3-4C respectively)
as measured by ΓSEA; a large toe load (L3∝A)
as measured by Γhard stop; or the extended presence of a large
ankle load (L3-4D) as measured by Γankle. That said,
to detect these conditions reliably, a "guard condition" may first be
applied to the logic to reject any such noise in ΓSEA and
Γhard stop that may arise during the swing phase. This may be
a result of the SEA torque applied to achieve the ankle trajectory and a
possible collision with the hardstop during the time interval. The LSW-ES
guard logic (GUARD) may be implemented as follows:

In the event that GUARD is FALSE, the LSW to ES state transition (3-4)
logic may be as follows:

L3-4=L3-4AORL3-4BORL3-4CORL3-4.s-
ub.D

where

L3-4A:(Γhard stop>45 Nm)AND

(Γhard stop(t)-Γhard stop(t-40 msec)>11 Nm).

L3-4B:(min(ΓSEAes)detected)AND

(Motor is in the READY state) AND

({dot over (Γ)}SEA<-50 Nm/s)AND

(ΓSEA<min(ΓSEAes)-2 Nm).

L3-4C:(min(ΓSEAes)detected)AND

({dot over (Γ)}SEA<-180 Nm/s)AND

( ΓSEA.sub.[t,t-6 msec]<min(ΓSEAes)-1
Nm)AND

(ΓSEA(t)-ΓSEA(t-6 msec)<-0.5 Nm)AND

(ΓSEA(t)-ΓSEA(t-10 msec)<-1.0 Nm).

L3-4B:(tLSW>1500 msec)AND

(TransitionEnabled=TRUE)AND

Γankle(t)>30 Nm)V t where tLSW-300
msec<t≦tLSW.

where tLSW is the elapsed time since LSW entry, ΓSEA(t),
and Γhard stop(t) are the SEA and Hard Stop torque at time, t,
respectively, READY is a signal indicating that the motor controller
processor has completed the trajectory return, Transition Enabled is a
motor state indicating that the motor controller has completed the
trajectory return instruction and that the motor temperature measurement
has been completed. AnkleNotReturned is a check to indicate whether the
ankle has returned to an initial state and has suitably dorsiflexed.
min(ΓSEAes) is the first validated minimum of SEA torque
while GUARD=FALSE. ΓSEA.sub.[t,t-n msec] is notation for the
mean of ΓSEA computed using samples from the prior n
milliseconds referenced to the current time, t.
Γankle(t)=ΓSEA(t)+Γhard stop(t) is the
total ankle torque.

[0099] For various embodiments presented herein, it is noted that the ES,
LS, LSP, ESW and LSW control response may be invariant with respect to
which logic condition--L3-4A, L3-4B, L3-4C
or L3-4D--causes the state transition into ES.

Late Stance (LS)-to-Early Stance (ES) Transition

[0100] In some cases, for instance, when the wearer stops in mid-stance,
the control system may transition from LS (State 5) back to ES (State 4),
so that the ankle state responds in accordance with the true walking
cycle state. The L5-4 transition may be informed by a negative
change in ΓSEA after the elapsed time in LS exceeds 500 msec
and may be summarized as follows:

where tLS is the elapsed time since entering LS
maxLS(ΓSEA(t)) is the maximum value of ΓSEA(t)
in LS.

Early Stance (ES)-to-Early Stance (ES) Transition

[0101] In some cases, for instance, when the wearer stops in ES then
begins to walk again, the impedance and equilibrium are reset to
appropriate values for foot strike to occur. Accordingly, the device may
be configured to re-enter the ES state based upon detection of an
L4-4 transition. This transition may be informed by a negative
change in ΓSEA after the elapsed time in ES exceeds 500 msec,
and may be summarized as follows:

where tES is the elapsed time since entering ES
maxES(64SEA(t)) is the maximum value of ΓSEA(t) in
ES.

Late Stance Power (LSP)-to-Early Stance (ES) Transition

[0102] In some cases, the entry into ES from LSP may occur if the ankle is
back-driven into LSP (LSPRegen)--to protect the wearer in the event that
the state machine does not detect a walking state transition out of LSP,
for example, to ESW. Because there is no stiffness in opposition to a
plantar flexion displacement in LSP, the expected ES impedance
(heel-strike stiffness) may be absent in a heel-strike event and would
thereby surprise the wearer. That is, if there is no stiffness in the
ankle after LSP occurs, the system may, by default, set its parameters to
the ES stance in preparation for the device in striking the ground.

LSP-to-ES "LSPRegen" Transition The LSP-ES LSPRegen transition may occur
when L6-4LSPRegen=TRUE per the logic equation:

where t=tLSP is the elapsed time in LSP and maxLSP
ΓSEA is the maximum value of the SEA torque since entry into
LSP.

Late Stance (LS) Control Response

[0103] In various embodiments of a controller for artificial leg devices
presented herein, LS (State 5) bridges the control response between ES
and LSP--typically between foot flat and hard stop engagement. In LS, the
actuator continues to apply a damped, torsional spring response so as to
correspond with the early CD response in human biomechanics.
Mathematically, the LS response is captured in Eq. 1.

[0104] It is well-understood that the spinal reflex arc connecting the
Achilles tendon stretch and the soleus (calf) muscle contraction employs
positive force feedback--both torque and torque derivative are employed
to amplify the reflex response in the contractile element (muscle). To
mimic this reflex arc in artificial leg devices according to the present
disclosure, the peak rate of change of ankle torque in LS, {dot over
(Γ)}anklels, may be used as input for the strain-rate
component of the reflex and spring dynamics applied in LSP by the
SEA--itself the bionic, artificial muscle-tendon unit in the BiOM ankle.
Here,

{dot over (Γ)}anklels={cSEAmaxls({dot over
(Γ)}SEA)+chsmaxls({dot over (Γ)}hard
stop)}ls (1)

[0105] In some embodiments, ES entry into LS (State 5) is the only state
transition into LS. The LS transition may occur if either a large toe
load (L4-5A) or heel load (L4-5B) is sensed by
Γhard stop and ΓSEA respectively. An example of the
mathematical formulation of the state transition (4-5) is described
below.

[0107] It should be appreciated that the control response in LS, LSP, ESW
and LSW may be invariant with respect to which logic
condition--L4-5A or L4-5B--causes the 4-5 transition.

Late Stance Power (LSP) Control Response

[0108] In some embodiments, the actuator response in LSP (State 6) is
comprised of two terms--a unidirectional torsional spring, klsp,
with equilibrium at a torque-rate-dependent plantar flexion angle,
θpp, and a torque-rate-dependent reflex. The reflex term
applies a positive force-feedback response that comprises two
components--a torque-rate dependent gain, pff({dot over
(Γ)}anklels), and a non-linear, normalized joint torque
feedback, {tilde over
(Γ)}ankle=ΓSEA+Γhard stop/Γ0,
with a torque-rate dependent exponent, N({dot over
(Γ)}anklels). In some embodiments, the torque gain may
range between 0 and 200 Nm, and the torque exponent may range between 1
and 5. Here, {dot over (Γ)}ls, is the peak rate of change of
joint torque in LS, as described in the previous section that addresses
late stance. Both pff and N may be piecewise-continuous, linear
functions, defined each by their values at a slow speed and a high speed
torque rate--{dot over (Γ)}anklelsslow and {dot over
(Γ)}anklelsfast respectively. At torque rates beyond
this range both pff and N may be held constant. In some embodiments,
pff and/or N are time-dependent functions, for example, that exhibit
an exponential decay behavior.

[0109] Mathematically, the LSP control response may be defined in Equation
2, shown below.

where τm is the SEA motor torque klspmax is a
torsional spring stiffness defined as the maximum of a quantity equal to
the torque-rate dependent reflex torque divided by the value of
θ-θpp. θpp is a plantar flexed torsional
spring equilibrium that is a piecewise, continuous linear function of
{dot over (Γ)}anklels pff and N are each a
piecewise, continuous linear function of {dot over
(Γ)}anklels as defined above, or may be time-dependent
functions that may range between 0-200 Nm and between 1-5, respectively.

Γ ~ ankle = Γ SEA + Γ hard stop
Γ 0 ##EQU00011##

is the normalized ankle torque where Γ0 is a normalizing
torque equal to

1.7 Nm kilogram m wearer ##EQU00012##

where mwearer is the wearer mass in kilograms.

[0110] In some cases, one or more of the parameters of an actuated torque
are time-dependent functions that exhibit time-decay behavior (e.g.,
exponential, linear, piecewise, etc.). For instance, kpp,
θpp, pff and/or N may exhibit exponential decay behavior
over time, so as to provide for a soft reflex response or gradual joint
equilibrium transitions. As an example, during LSP, the wearer may decide
that he/she would like to ease in or out of powered plantar flexion. If
the gain and/or exponent of the torque reflex response exhibits
time-dependent decay, the wearer may experience a relatively smooth
reflex response than may otherwise be the case without the decay
behavior. Or, θpp may also exhibit time-dependent decay
behavior, resulting in relatively smooth transitions from one state to
another. Any suitable time-dependent behavior may be employed, such as
those functions described for various embodiments of the present
disclosure. FIGS. 24-25 show examples of suitable reflex parameter
modulation relationships.

Late Stance Power (LSP) Entry State-Transition Details

[0111] In some embodiments, the LS to LSP transition (5-6) may occur when
the toe-load torque exceeds a programmable threshold. Mathematically, the
L5-6 transition may occur when Γhard stop>5 Nm.

Early Swing Control Response

[0112] In some embodiments, the ESW (State 2) control response of the
artificial leg device mimics the damped, second-order, spring-mass
response of the early swing phase in human walking biomechanics--this
response restores the ankle from the toe-off position at the terminus of
powered plantar flexion to its neutral position, in anticipation of the
foot strike in the next gait cycle. Typically, the time constant,
τesw, of this response is approximately 50 milliseconds, but may
vary appropriately.

[0113] In ESW, an overdamped, second-order equilibrium trajectory,
θ0(t), may be applied to return the joint to a fixed neutral
position, θesw--a position that may be invariant to all
biomechanical modalities including, but not limited to, terrain, walking
speed, and toe-off angle. A damped (besw) and spring (kesw)
impedance may be applied in relation to this equilibrium trajectory.
Feedforward of the estimated motor torque may be used to eliminate
response lag due to motor/drive-train inertia and damping. The
mathematical formulation of the ESW control response with inertia-only
feedforward may be summarized in Equation 3 shown below.

where τm is the SEA motor torque, τesw is the time
constant of the over-damped, second-order response, θ0(0) is
the toe-off angle, initialized to θ(t) at ESW entry (LSP Exit),
β is the SEA motor angle reflected at the ankle joint and
θesw is the invariant, neutral position destination for all
ESW trajectories J.sub.βm is the motor inertia reflected onto
the joint

Early Swing (ESW) Entry State Transition Logic

[0114] Transitions into ESW may normally originate from LSP, as described
in the following section that addresses the late stance power to early
swing transition. Transitions into ESW can originate from ES when the
wearer lifts the foot off the ground, as described in the section that
addresses ES-to-ESW at Foot-off.

Late Stance Power (LSP)-to-Early Swing (ESW) Transition

[0115] The LSP-ESW transition may be defined by either a toe-off
(L6-2toe-off) or a foot-off event (L6-2foot-off)
while in LSP.

LSP-to-ESW at Toe-Off

[0116] Toe-off may occur when the ankle torque, Γankle, drops
below a threshold close to zero.

[0117] The following guard, pre-trigger, and state transition conditions
may be applied in succession to accomplish the LSP-ESW (6-2) transition
by toe-off.

Toe-Off Guard Condition Details

[0118] The LSP-ESW by toe-off transition may be halted until GUARD has
transitioned from TRUE to FALSE.

where in the above, tlsp is the time since LSP entry
maxlsp(Γhard stop(t)) is the maximum value of hard stop
torque in LSP prior to tlsp {dot over ( ΓSEA.sub.[t,t-10
msec] is the mean value of SEA torque rate over the past 10 milliseconds.
As a result, the LSP to ESW transition can occur when L6-2 is TRUE.

LSP-to-ESW at Foot-Off

[0121] The "foot-off" condition--L6-2foot-off--may be informed
by a rapid drop in both SEA and Hard Stop torque, which may be summarized
as follows:

where t=tES is the elapsed time since entry into ES, FromLSPRegen is
a flag set in ES to note that ES entry originated from LSP during an
unexpected regeneration event in powered plantar flexion,
Guardfoot-off is a guard logic condition that blocks the transition
if ES entry originated from the excessive regeneration event in LSP or if
the elapsed time within ES is less than a pre-specified duration (800
milliseconds).

Late Swing (LSW) Control Response

[0123] In LSW after the ESW return to the neutral angle is completed, the
SEA applies a lightly-damped, torsional spring response equivalent to
that applied at ES entry. This ensures that the intended impedance to be
applied at foot strike is instantiated before impact-thereby achieving
response continuity that is insensitive to ES state transition delay. The
mathematical formulation of the LSW response is captured in Equation 4.

where θes0=θ(0) β is the motor angle as
projected onto the joint angle from SEA kinematics

[0124] In LSW, after the ESW return to the neutral angle is completed, the
SEA may apply a lightly damped, torsional spring response--with a spring
constant, kes(t) that may be designed to decay exponentially,
according to a second-order differential equation. Such a decay, while
not limited to exponential behavior, may help to ensure that the intended
impedance to be applied at foot strike is instantiated before
impact--thereby achieving foot-strike response continuity that is
insensitive to ES state transition delay. Such a form of decay dynamics
has the emergent property that stiffness decreases with increased walking
speed. This property acts to reduce foot-strike stiffness while walking
slowly down a steep slope, for instance. The joint equilibrium,
θes0, may be set to the ankle angle, at entry, θ(0). The
mathematical formulation of the LSW response, including stiffness decay
dynamics, is captured in the Equations 5 and 6 below.

[0129] β is the motor angle as
projected onto the joint angle from SEA kinematics

[0130]
τkes controls the stiffness decay, typically 200
milliseconds

[0131] kes(0)=kes0

[0132]
kes.sub.∞ is the terminal value of stiffness

Late Swing (LSW) Entry State-Transition Details

[0133] The ESW-LSW state transition may occur when the motor control
processor reports that it is READY, thereby signifying that the ESW
trajectory is completed, OR, for example, when tesw>100 msec.

Late Swing (LSW) Entry from Early Stance (ES)

[0134] An ES-LSW transition can occur in cases where after an extended
period in ES (e.g., approximately two seconds) a possible ground impact
is present as detected by a toe load (L3-4A), toe unload
(L3-4B), or footstrike (L3-4C), as provided below.

[0135] L4-3A: Toe-Load Detected

[0136] (Γhard
stop>45 Nm) AND

[0137] (Γhard stop(t)-Γhard
stop(t-40 msec)>11 Nm).

[0138] L4-3B: Toe-unload Detected

[0139]
(min(ΓSEAes) detected) AND

[0140] (Motor is in the READY
state) AND

[0141] ({dot over (Γ)}SEA<-50 Nm/s) AND

[0142]
(ΓSEA<min(ΓSEAes)-2 Nm).

[0143] L4-3C: Foot-Strike Detected

[0144]
(min(ΓSEAes) detected) AND

[0145] ({dot over
(Γ)}<-180 Nm/s) AND

[0146] ({tilde over
(Γ)}SEA.sub.[t,t-6 msec]<min(ΓSEAes)-1 Nm)
AND

[0147] (ΓSEA(tes)-ΓSEA(t-6 msec)<-0.5
Nm) AND

[0148] (ΓSEA(t)-ΓSEA(t-10 msec)<-1.0
Nm).

[0149] Where

[0150] tes is the elapsed time since ES entry,

[0151] ΓSEA(t), and Γhard stop(t) are the SEA and
hard stop torque at time, t, respectively,

[0152] READY is a motor state
indicating that the motor controller processor is ready to accept
commands.

[0153] min(ΓSEAes) is the first validated
minimum of SEA torque after ES entry.

[0154] While description for each of the state transitions is provided
above, Table 3 summarizes the state transition logic, including various
non-limiting conditions and thresholds that are used for an embodiment of
an artificial leg device, in accordance with the present disclosure. FIG.
11 provides a schematic that illustrates operation of an embodiment of an
artificial leg device.

[0155] Embodiments of the present disclosure may include a multi-modal
control system for an artificial leg device having series and
parallelelastic actuator-based muscle-tendon units (MTU) at the ankle and
knee for modulation of joint impedance, joint equilibrium and reflex
torque, in accordance with locomotion modality, gait cycle phase within
that modality and cadence; a plurality of metasensors for intra-gait
cycle determination of terrain modality, ground reaction force and
zero-moment point, and external load-bearing influence; an intent
recognition processor that employs the metasensor data to infer
locomotion modality and the transitions between these; and a
biophysically-inspired state control processor that employs MTU torque
and derivatives, metasensor state and intent recognition output to
accomplish transitions between the joint-based state machines.

[0156] The bionic architecture may restore function per normative measures
of metabolic cost-of-transport and gait mechanics, including joint
kinematics and kinetic measures. The architecture may further optimize
battery economy and achieve safe operation in the event of power loss
through use of tuned series-elastic elements and regenerative dynamic
clutching (braking) functions in the joint MTU controls.

[0158] FIG. 12 illustrates elements of another embodiment of a bionic leg
system architecture in accordance with the present disclosure. In the
embodiment shown, the system includes series-elastic actuators (SEA)
serving as bionic muscle-tendon units (MTU) at the ankle and knee; an
ankle socket-mounted force/torque sensor to measure axial force, sagittal
plane moment and coronal plane moment; a state control processor that
embodies a gait cycle state machine, modulates MTU response, and
recognizes wearer intent-including terrain (sloping ground and stairs)
context. Here, intent recognition can be accomplished through use of
metasensors as follows:

Kinematic State Estimator (KSR)--

[0159] The KSR employs a 6-DOF IMU embedded in the ankle or knee and the
knee joint angle, θk to reconstruct the tibia and femur
coordinate systems in real-time-capturing the inertial path of the ankle,
knee and hip and points between these throughout all or part of a gait
cycle.

Terrain Modality Discriminator (TMD)--

[0160] The TMD applies pattern recognition of the ankle, knee and hip
translational and rotational paths during the swing phase to infer
underlying terrain. The state control processor uses the terrain context
to inform the ankle and knee equilibrium and impedance at foot strike.

Ground Reaction Force/ZMP Estimator (GRFZMP)--

[0161] The GRFZMP processes the force-torque sensor data, the ankle joint
torque and the tibia kinematic state to compute the ground reaction force
vector and the zero-moment position of this. This information may be used
by the state control processor in combination with the KSR, TMD and EIE
(below) to determine locomotion context (walking, sitting, standing,
stair climbing) and/or to apply balance control while standing, walking
and running.

External Influence Estimator (EIE)--

[0162] The EIE may use the GRFZMP and the KSR information to determine,
via inverse dynamic approximation, the external influences that must be
acting on the trunk (as measured at the hip) to achieve its kinematic
state (of acceleration). The EIE can estimate, for instance, the
presence, and influence of external force as might be applied by the arms
as the bionic leg wearer lifts out of a chair. The EIE can also estimate
the presence and influence of trailing leg powered plantar flexion on a
stair. Such information may be used by the state control processor to
determine when to apply leg joint torques in such locomotion contexts.
Additional details regarding various embodiments of the leg architecture
are provided in the references incorporated by reference above.

Control Architecture

[0163] Embodiments of the leg system employ a loosely-coupled joint
control architecture. Here, the ankle state machine and control behaviors
are largely independent of the knee control state. Ankle state machine
and control behaviors are described in greater detail in the references
incorporated by reference above. In particular, the
biophysically-motivated ankle state machine and behaviors are described
in detail in U.S. Provisional Patent Application Ser. No. 61/662,104,
entitled "Bionic Control System for an Artificial Ankle Joint."

[0164] A schematic of one embodiment of a knee state machine is
illustrated in FIG. 13. As shown in FIG. 13, the Knee State Machine (KSM)
embodies four states-Early Stance, Late Stance, Swing Flexion and Swing
Extension with state-dependent control behaviors and state transitions
(ST1, ST4, ST6, ST7 and ST8), as further discussed below.

State-Dependent Control Behaviors

[0165] FIG. 14 illustrates the kinematic behavior of the knee during a
typical gait cycle where ES refers to Early Stance; LS refers to Late
Stance; ESW refers to Early Swing; LSW refers to Late Swing;
θk refers to Knee Angle; HS refers to Foot Strike; and TO
refers to Toe-Off.

Early Stance

[0166] In Early Stance, the knee applies a lightly-damped spring response
defined by stiffness, kES and damping, bES0. For stance
flexion, δθk=θk-θ0es, when less
than about 15°, the early stance impedance relation may be
provided as follows:

[0169] Equations 7 and 8 may be implemented by using closed-loop torque
control, using SEA deflection as a measure of joint torque feedback. In
another embodiment, the knee SEA may employ a series elasticity with
stiffness substantially equal to kes. In this way, the motor drive
transmission can be locked at θes0, enabling the series
elastic element to compress and extend without motor movement to account
for the maximum early stance knee flexion for typical level-ground gait
cycles.

[0170] In another embodiment, the motor may be employed as a programmable
clutch (dynamic brake/damper) by shorting the motor leads--applying a
strong braking function with a time constant typically in the range of
approximately 800-1500 milliseconds. Details concerning the use of
shorted leads may be found in U.S. patent application Ser. No.
13/417,949, entitled "Biomimetic Joint Actuators." In such an embodiment,
the battery power source may be disconnected from the SEA, thereby
eliminating battery consumption during knee flexion and extension in
level-ground walking.

[0171] In some cases, the shorted-leads may be pulse-width modulated,
enabling the damping to be controlled, e.g., to reduce the damping at
large flexion while at the same time harvesting energy to charge the
bionic leg power source (i.e., battery) during, for example, the swing
phase of walking. Since the knee joint generally draws net energy, such
an embodiment can be used to operate the knee joint at extremely low
power in at least early stance flexion/extension early swing and late
swing, even when the battery is disconnected. The shorted leads
functionality can make possible assertion of a safe state during fault or
power interruption, thereby protecting the wearer. In some embodiments,
kes, and bes, are functions of time (e.g., may exhibit a
time-dependent decay behavior). For instance, the change from a
stiffness-dominated response to the damping-dominated response may not be
accomplished by crossing an angle threshold, but rather by applying a
programmable, exponential decay of the stiffness and damping as shown in
FIG. 15, which illustrates an early stance exponential stiffness and
damping response for an embodiment of a knee device.

[0172] The stiffness and damping impedance coefficients may be defined by
the following relations:

32 τk2{umlaut over (k)}es(t)+2τk{dot over
(k)}es(t)+1=kesmin Eq (9)

Where

[0173] kes(0)=kesmax and τk is the time constant
of the stiffness decay

τb2{umlaut over (b)}es(t)+2τb{dot over
(b)}es(t)+1=besmin Eq (10)

Where

[0174] bes(0)=besmax and τb is the time constant
of the damping decay

[0175] As shown in FIG. 15, the time constant for stiffness decay may be
set to be shorter than the damping time constant. Though, in some
embodiments, the time constant for stiffness decay may be greater than
the damping time constant.

[0176] In some embodiments, a first-order or higher order differential
equation may be used in place of Eqs. 9 and 10. A second-order response
may be advantageous in that the attenuation is substantively delayed--the
initial values are substantially maintained for a certain amount of time
controlled by the time constant prior to dropping off. Through these time
varying impedances, the knee will behave during early stance as an
efficient spring during level ground walking, a damper with a relatively
high damping value for stair and slope descent, and a lightly damped knee
while sitting.

Late Stance

[0177] The joint torque sign reversal at substantially full knee extension
signals the transition from Early Stance to Late Stance in a typical gait
cycle. In one embodiment, the Late Stance reflex behavior follows the
relation below:

[0178] τmotorknee is the SEA motor torque.
Γ0k is a normalizing torque defined by body weight and
activity level, and pff( ) and N( ) are functions of knee torque rate of
change at entry to late stance.

[0179] In other embodiments, a neuromuscular model, also employing
positive force feedback on a modeled Gastrocnemius muscle, may be used.
For further details regarding this neuromuscular model, the disclosure of
U.S. Provisional Patent Application Ser. No. 61/595,453, entitled
"Powered Ankle Device" may be relevant.

[0180] In certain cases--including stair ascent, steep ramp ascent and
during the transition from sitting to standing--the knee joint may be
flexed past a threshold of θkts0 and extending at a
substantial rate (|{dot over (θ)}k|>{dot over
(ξ)}ext0) where {dot over (ξ)}ext0 is the rate
threshold. In this case, a rate dependent spring stiffness, kex,
that applies positive feedback in response to angular rate increases for
an embodiment of a knee device as shown in FIG. 16a and captured in the
anti-slip impedance control behavior defined by Eq. 12 may be applied.

Γk=-kes({dot over ( θk)
θk-bex( θ) θk Eq (12)

Where

[0181] bex( θ) applies light damping to achieve stability when
{dot over (θ)}≦0, bex( θ) applies strong damping
to resist flexing θk is the knee joint angular rate, and
θ is the output of a peak detection filter of the form

τext({dot over (θ)})+ θ+ θ={dot over
(θ)}, where

τext( θ)=τextsmallif θ<0and

τext( θ)=τextlarge if θ≧0

And where kex({dot over ( θ) is of the form shown in FIG. 16a
In some embodiments, kex and bex, are time-dependent functions
that exponentially decay over time and are initialized to the nominal
form when retriggered ( θ≦ξext0). In an
"anti-slip" embodiment described here, momentary flexion velocities do
not cause the knee torque to drop-thereby making it easier for the wearer
to maneuver (e.g., to get out of a chair or to transition to bionic limb
support when the sound side (trailing leg) is pushing off of a stair
below the bionic limb). FIG. 16s defines the general form of kex,
illustrating that the flexion stiffness may increase with increasing
joint speed. In some cases, the peak flexion stiffness may have a lower
peak than in extension, enabling the wearer to more easily flex the knee
while sitting. FIGS. 16b-16d illustrate schematics of a wearer moving
from a sitting position to a standing upright position.

Swing Flexion

[0182] The Early Swing state transition occurs at toe-off, as reported by
the ankle state machine. In early swing flexion, knee behavior may be
ballistic for flexion angles less than about 45° (e.g., no spring
or damping) and lightly damped (b=bsf) for greater flexion. This
behavior is captured in Eq. 13.

Γ k = - b sf θ . k θ k
< θ sf = 0 elsewhere Eq . (
13 ) ##EQU00018##

Swing Extension

[0183] Once the maximum swing flexion is achieved, the knee state
transitions to swing extension. In early swing extension the behavior is
nearly ballistic (e.g., lightly damped) with damping constant,
bse=bse.sub.∞. The damping coefficient increases nearly
quadratically as the knee flexion approaches
θk=θkes0, as shown in three piecewise
continuous angle-dependent damping function embodiments (in swing
extension) in FIGS. 17a-c. FIG. 17a depicts the behavior of an embodiment
that exhibits piece-wise constant and linear behavior. FIG. 17b
illustrates the behavior of an embodiment that exhibits piece-wise linear
and quadratic behavior. FIG. 17c shows the behavior of another embodiment
that exhibits a more general functional form.

[0184] In Swing Extension, such behavior may be captured in Eq. 14.

Γk=-bse(θk) θk Eq. (14)

Where

[0185] bse(θk) is defined as a piecewise continuos
function per FIGS. 17 a-c

[0186] Damping during Swing Extension may be used to decelerate knee
flexion (tibia angular rate) as the joint angle approaches
full-extension--increasing substantially linearly until θ drops
below a threshold angle. Below the threshold, the damping increases
according to a substantially quadratic function as it approaches
θ≈0. Such damping creates a "sticky" behavior that holds
the joint near full-extension-preparing the knee to absorb the foot
strike energy and to transition to the spring-like behavior in Early
Stance.

State Transitions

[0187] FIG. 13 illustrates the knee state machine and defines knee
controller state transitions, as further discussed below.

[0188] The foot strike gait event marks the transition from Swing
Extension (or Flexion)-to-Early Stance--a transition that aligns with the
Late Swing to Early Stance transition on the ankle. Here, the world-z
component of the ground reaction force, as shown in FIG. 18, will be used
to detect the ST1 transition (i.e., foot strike transition, heel-strike
or toe-down), defined as:

ST1=(FZ>FZFS) Eq. (15)

Where FZFS is the force transition threshold that signals
foot-strike.

[0189] In another embodiment as described in U.S. Provisional Patent
Application Ser. No. 61/662,104, entitled "Bionic Control System for an
Artificial Joint," a logic transition informed by ankle torque and
derivatives can be used to accomplish ST1.

State Transition 6 (ST6): Early Stance-to-Late Stance

[0190] The Early Stance to Late Stance transition gait event signifies
that toe-loading is occurring when the knee is fully extended as defined
by the logic equation:

[0191] ξ.sup.+ and ξ.sup.- are small angles signifying proximity
to full extension, and Γtoe load is the toe loading threshold
as measured at the ankle, and Γa signifies the ankle torque
reported by the ankle MTU.

[0192] In other embodiments, toe loading is detected by determining
whether the ZMP of a ground reaction force of significant magnitude is
substantially located in the forward half of the foot.

State Transition 4 (ST4): Late Stance (or Early Stance)-to-Swing Flexion

[0193] The toe-off gait event signals the transition to Swing Flexion from
either Late Stance or Early Stance. ST4 is defined as:

ST4=(FZ<FZtoe off) Eq. (17)

Where FZ is the z-component of the ground reaction force, and
FZtoe off is the toe-off force threshold. In other embodiments,
substantially zero torque, as reported by the ankle MTU, can be used to
detect the toe-off condition. In another embodiment described in U.S.
Provisional Patent Application Ser. No. 61/662,104, entitled "Bionic
Control System for an Artificial Joint," ankle torque and derivatives
(Γankle≈0) can be used as input for triggering or
modulating parameters of the ST4 transition.

State Transition 7 (ST7): Swing Flexion-to-Swing Extension

[0194] The state transition from Swing-Flexion to Swing Extension is
marked by a sign reversal in the knee angular velocity-detected here as
the time when the knee velocity goes to zero at a time sufficiently after
toe-off:

Where tsf is the time elapsed since toe-off, tsfmin is
the minimum duration threshold, and {dot over (ξ)}.sup.+ and {dot over
(ξ)}.sup.- define the small velocity boundary.

State Transition 8 (ST8): Late Stance-to-Early Stance

[0195] In some circumstances, e.g. when the wearer is standing quietly and
then enters Late Stance and then flexes the knee, it may be appropriate
for the state machine to transition back to early stance. The logic is
defined as follows:

[0196] θklarge defines the angle threshold, and
ξΓ.sup.- and ξΓ.sup.+ define the small torque detection
boundaries.

Other Embodiments

Self-Adjusting Joint Equilibrium

[0197] In this embodiment, the joint equilibrium tracks the joint angle
with a programmable convergence--preferably through use of a first or
second-order tracking filter with time constant τ. In some
embodiments, the system is configured for the joint equilibrium to
exhibit time-dependent behavior that relaxes to an equilibrium that is
substantially equivalent to the current joint angle. That is, in
accordance with the system exhibiting a programmable convergence, the
joint equilibrium of the system continually, yet gradually, tracks the
current joint angle. For example, if the joint angle does not change
after a long period of time, then the joint equilibrium gradually relaxes
from an initial value to a value equal to that of the current joint
angle.

[0198] In some embodiments, self-adjusting joint equilibrium behavior may
be governed by the following relationships:

Γ=-k(θ-θ0)-b{dot over (θ)} Eq. 20

τ0{dot over (θ)}0+θ0=θ Eq. 21

Equation 21 is inserted into Eq. 20 and the resulting relationship is
subject to a Fourier transform, where the function is transformed from
the time domain to the frequency domain. Accordingly, the derivative
represented by ({dot over ( )}) is replaced with s=jω and
ω0 with 1/τ74 resulting in an impedance relation of
the form:

Γ ( s ) = { ( b + K ω 0 ) H ( s )
} s θ ( s ) Eq . 22 ##EQU00022##

where H(s) is defined by the relation,

H ( s ) = ( s k b + ω θ + 1 s
ω θ + 1 ) Eq . 23 ##EQU00023##

[0199] FIG. 19 illustrates the impedance transfer function,

Γ ( s ) θ ( s ) , ##EQU00024##

represented by Eq. 22.

[0200] The frequency response of this impedance law has interesting
properties. At low frequencies, the impedance behaves as a damper with
coefficient, b*=b+kτ. At medium frequencies, the impedance has
stiffness properties with an equivalent stiffness of

k * = k + b τ . ##EQU00025##

And at high frequencies, the impedance behaves as a damper with
equivalent damping of

b * = ( b + k τ ) ( ω θ ω
θ + k b ) ##EQU00026##

where

ω θ = 1 τ ##EQU00027##

is the transition frequency between the first damping and stiffness
behaviors. Here, ω74 may range from 0-13 rad/sec (0-2 Hz)
providing a primarily damping response in that range. Between Wtheta
and about 60 rad/sec (a preferred range between 5-20 Hz), a stiffness
dominated response is applied. Above this latter frequency defined by

ω θ + k b , ##EQU00028##

damping-dominated response is applied. Often wearers complain that it is
hard to maintain balance when the leg joints are in a substantially
lightly damped state. So by implementing this method, improved stability
results because in the frequency range between 1-10 Hz a
stiffness-dominated response is applied that serves to restore balance.

Blended Reflex

[0201] The following disclosure describes two blended reflex methods, each
blending (interpolating) independently tuned responses--defined by torque
gain (Pff) and torque exponent (N), at a fast and a slow walk speed.
At speeds below the "slow-walk" speed as determined by the wearer (e.g.,
less than 0.75 m/s), the reflex employs a slow-walk parameter set; at
speeds greater than the fast-walk speed as determined by the wearer
(e.g., greater than 1.75 m/s), the reflex employs the fast walk parameter
set; and at speeds in between, the reflex adds the two responses together
in accordance with a linear or non-linear interpolation based upon
walking speed, a surrogate for walking speed (e.g., pitch rate in
mid-stance), a kinetic (e.g., torque rate) or kinematic (e.g., joint
angle rate). The term walking speed and operating speed below may loosely
refer to the walking speed, surrogates of walking speed, a suitable
kinetic rate or a suitable kinematic rate.

[0202] Other interpolations may be used, and more than two
speed-registered responses may be blended through more complex
interpolation, for example, based upon the "distance" between the
operating speed and each of the tuned speeds. This approach may be
advantageous over the existing methods in that both the gain and exponent
can be independently controlled--that is, these reflex coefficients can
be tuned independently of each other. For instance, a slow walk reflex
response may require a lower exponent torque than that required by a fast
walk reflex response, and vice-versa. With fixed N (the variable that
controls timing), there is a tradeoff between slow-walk consistency and
fast walk power and battery economy. By applying independent tuning, an
optimum performance may be achieved at both ends of the walking speed
spectrum, and overall wearer experience can be improved.

[0203] Method I blends two torque models--one defined at a slow speed and
one at the fast speed, as determined by the wearer--with gain,
Pff({dot over (s)}slow), and exponent, N({dot over
(s)}slow), for a first ("slow-walk") torque model; and gain,
Pff({dot over (s)}fast), and exponent, N({dot over
(s)}fast), for a second ("fast-walk") torque model. Method II blends
the gains and exponents into a single torque model--with gain,
Pff({dot over (s)}), and exponent, N({dot over (s)}), where the gain
and exponent are speed interpolated (via linear or non-linear
interpolation) across the speed domain, [{dot over (s)}slow, {dot
over (s)}fast]. The blended torque models are expressed by suitable
computations below.

[0207] FIGS. 20-23 illustrate ankle data gathered from test subjects of
walking information that are used as design parameters for the control of
artificial leg devices in accordance with the present disclosure.
Accordingly, embodiments provided herein may employ this data to create a
dashboard of normative measures across walking speed that capture the
kinetics and kinematics of natural limbs. In some embodiments of the
control architecture described above, the kinetic and kinematic response
of the bionic ankle joint is projected onto this dashboard of normative
measures. The impedance, equilibrium and torque, including reflex,
modulation may then be optimized to fit within the normative statistical
range noted in the dashboard. Bionic restoration of ankle-foot function,
as measured by the closeness of fit, is thereby achieved. And this
projection of kinetic and kinematic measures onto the dashboard serves as
a record that can be used by the clinician to prove the efficacy of the
bionic limb as this might be needed for insurance reimbursement or other
purposes.

[0208] FIG. 20 shows graphs that depict ankle angle, angular velocity,
moment, and power plotted as a percentage of the gait cycle. Plots are
shown for the average of all subjects walking at their fast walking speed
(e.g., between 1.5-2.5 m/s). As further shown, the stance phase is
divided into three subphases--controlled plantar flexion (CP), controlled
dorsiflexion (CD) and powered plantarflexion (PP). Various embodiments of
the present disclosure may employ principles described in the Masters
Thesis by Gates, D. H, entitled "Characterizing Ankle Function During
Stair Ascent, Descent, and Level Walking for Ankle Prosthesis and
Orthosis Design," submitted in 2004, the disclosure of which is hereby
incorporated herein by reference in its entirety.

[0209] FIG. 21 depicts a scatter plot graph of the net non-conservative
ankle work (WNET=WCP+WCD)) on level-ground performed by
walkers with an intact ankle (population N=70) during the stance phase of
gait as a function of walking speed. Each point represents the average
work done for all trials of a subject when asked to walk at a certain
speed (fast, normal, slow). A linear regression was performed on the mean
work for each subject walking at his or her mean speed. This line shows a
significant increase in ankle work and linear correlation with gait
speed. The rate-dependent, blended reflex disclosed above may be
optimized to achieve a close fit to this linear net non-conservative
ankle work vs. walking speed relationship.

[0210] FIG. 22 illustrates the correlation between ankle torque and each
of ankle angle and ankle velocity, during a single gait cycle. Data are
shown for an average of all subjects walking at fast, normal, and slow
speeds. Trials were normalized to 50 equally spaced data points, which
was then averaged for each subject. Numbers mark the beginnings and ends
of subphases of gait (CP: 1-2, CD: 2-3, PP: 3-4). As shown, at normal
walking speeds, the ankle torque correlates strongly with ankle position
during these subphases. As further shown, the faster the walking speed,
the greater the net amount of work performed by the ankle (shown by the
area under the curve for the ankle angle versus ankle torque graphs).

[0211] FIG. 23 shows graphs of ankle torque versus ankle angle plotted for
each subphase of stance for walking subjects. Data are shown for the
average of all subjects walking at their self-selected slow, normal and
fast speeds. For the CP phase (top), there is a generally linear
relationship at each walking speed. For the CD phase (middle), the
relationship increases in non-linearity as speed increases. For the final
phase, PP, the fitting is generally linear.

[0212] It should also be understood that, unless clearly indicated to the
contrary, in any methods claimed herein that include more than one step
or act, the order of the steps or acts of the method is not necessarily
limited to the order in which the steps or acts of the method are
recited.

[0213] While aspects of the invention have been described with reference
to various illustrative embodiments, such aspects are not limited to the
embodiments described. Thus, it is evident that many alternatives,
modifications, and variations of the embodiments described will be
apparent to those skilled in the art. Accordingly, embodiments as set
forth herein are intended to be illustrative, not limiting. Various
changes may be made without departing from the spirit of aspects of the
invention.